In nuclear emission tomography, gamma cameras or detectors typically are used for locating and displaying human glands and organs and associated abnormalities. Abnormalities may be represented by higher uptake or lower uptake than the surrounding tissue. More specifically, and with respect to using a gamma camera, gamma-ray-emitting tracer material is administered to a patient, and the tracer material is more greatly absorbed by the organ of interest than by the other tissues. The gamma camera generates data, or an image, representing the distribution of such tracer material within the patient.
A gamma camera includes a multi-channel collimator and a gamma ray detector which converts energy from the gamma ray into an electrical signal which can be interpreted to locate the position of the gamma ray interaction in the planar detector. One known gamma ray detector which is commonly used is an Anger gamma camera, which is described in H. O. Anger, "Scintillation Camera", Rev. Sci. Instrun., Vol. 29, p. 159 (1958). Another known detector is a multi-crystal scintillation detector which has an array of small crystals coupled to an array of light detectors, which may be either photomultipliers or photodiodes. Yet another known detector is a solid-state position sensitive detector which converts energy from the gamma ray into an electrical charge which can be detected by an array of contacts.
The Anger gamma camera includes a large scintillation crystal responsive to radiation stimuli, i.e., gamma rays emitted by the patient. An array of photomultiplier tubes typically are optically coupled to the crystal. In operation, the gamma rays emitted by the patient in the direction of the detector are collimated onto the crystal, and each gamma ray which interacts with the crystal produces multiple light events. The multiple light events are detected by photomultipliers adjacent to the point of interaction. The photomultiplier tubes, in response to the light events, produce individual electrical outputs. The signals from the array of photomultipliers are combined using analog and digital circuitry to provide an estimate of the location of the gamma ray event. Further analog and digital processing is used to produce more accurate position coordinates to form the acquired image.
More particularly, to generate an image, a representation of the distribution of events in the crystal is generated by utilizing a matrix of storage registers whose elements are in one-to-one correspondence with elemental areas of the crystal. The crystal elemental areas are identified by coordinates. Each time a light event occurs in the crystal, the event coordinates are identified and the register in the storage register matrix corresponding to the identified event coordinates is incremented. The contents of a given register in the matrix is a number that represents the number of events that have occurred within a predetermined period of time within an elemental area of the crystal. Such number is directly proportional to the intensity of radiation emitted from an elemental area of the radiation field. The number stored in the register therefore is used to establish the brightness of a display picture element corresponding to the crystal elemental area. The distribution of a radiation field is displayed in terms of the brightness distribution of the display.
In emission tomography a plurality of such images are taken at various view angles around the organ of interest. Typically, in transaxial tomography, a series of images, or views, are taken at equal angular increments around the patient. The series of views around the patient are reconstructed to form transaxial slices, that is, slices across the axis of rotation. The process of acquiring the views and reconstructing the transaxial slices is termed emission computed tomography (ECT) or single photon emission computed tomography (SPECT). Similar reconstruction concepts are employed in X-ray computed tomography (CT) in which X-rays are used to measure patient attenuation and to reconstruct the attenuation in transverse sections.
Photon attenuation and scatter directly affect image resolution. Particularly, photon attenuation and scatter typically cause image artifacts and degrade SPECT image quality both qualitatively and quantitatively. One known method of reducing artifacts caused by photon attenuation is to utilize iterative reconstructions, such as maximum likelihood expectation maximization, and an attenuation map. While such method reduces photon attenuation artifacts, such method does not noticeably reduce artifacts caused by photon scatter. Images generated with only photon attenuation correction, i.e., without also correcting photon scatter, often demonstrate reduced contrast, or even artificially increased or decreased activity levels.
Many methods have been implemented to compensate for the detection of scatter photons such as Compton scatter photons. Such methods include utilizing deconvolution techniques, spectral fittings methods, and matrix-based inverse Monte Carlo techniques. Each of these methods, however, is complex and substantially difficult to implement in a SPECT system.
Another method for compensating for scatter photons, which is more easily implemented, is known as the dual energy window scatter subtraction method. In this method, a first window is placed over the primary energy window, or photopeak window, and a second window is placed below the photopeak window. Projections acquired from the second energy window are multiplied to a single-value scatter fraction, which typically is either object independent or source location-dependent, to generate an estimated scatter projection in the primary energy window. The estimated scatter projection in the primary energy window is then subtracted from the projections acquired from the first window to estimate the non-scatter photons in the primary energy window. Accordingly, the dependence of the scatter component of the photopeak image on surrounding activity is determined on a patient specific basis, reducing the number of assumptions required in the correction methodology. The known dual energy window scatter subtraction method, while generally successful, requires significant operator intervention to determine the single-value scatter fraction.
It would be desirable to provide more accurate scatter correction and reduce operator intervention necessary to achieve such improved correction. It also would be desirable to provide a patient-dependent scatter fraction, as compared to a single-value scatter fraction.